1. Field of the Invention
The present invention relates to improvements in ultrasonic imaging, and more particularly, to a novel connection arrangement and method of operation of a 2D array which improves the performance of phase aberration correction processing while minimizing increases in circuit and processor complexity. The array includes first and second portions which are segmented into differently dimensioned elements in their elevation direction, wherein the first portion of the array is segmented into coarse elements and the second portion is segmented into fine elements. The connection arrangement selectively connects a plurality of the fine elements together so that the area of the 2D array corresponding to the connected fine elements is substantially the same size as the area of a coarse element.
2. Description of the Prior Art
Ultrasonic imaging has been extensively applied in virtually every medical specialty in the form of pulse-echo (B-mode) imaging. B-mode imaging systems display echoes returning to an ultrasonic transducer as brightness levels proportional to the echo amplitude. The display of brightness levels results in cross-sectional images of an object in a plane perpendicular to the transducer.
Current ultrasonic transducer arrays typically include transmit mode focusing and receive mode dynamic focusing, achieved by appropriate timing of the transmit signals and appropriate delaying of the received echoes. For example, a linear phased array consists of a single group of transducer elements arranged in a line which are operated to not only focus but also steer (angle) transmit and receive beams by appropriate timing of the transmit signals and the receive echoes.
Conventionally, the timing or phasing data is determined by assuming propagation of the ultrasound pulses through a homogeneous tissue medium with a uniform velocity of sound, usually 1540 m/sec.
The assumption of a constant velocity of sound in the body is also the design basis of all ultrasound scanning systems for converting round trip pulse-echo time of flight data into images. Unfortunately, this simple model for all human tissue is erroneous. The body is actually composed of a plurality of inhomogeneous layers of different tissues (fat, muscle and bone) with bumps and ridges of varying thicknesses and shapes, and therefore different acoustic velocities. These layers are situated between the transducer and, for example, an internal organ of interest. The propagation velocity of ultrasound varies from approximately 1470 m/see in fat to greater than 1600 m/sec in muscle and in nervous tissue, and to as much as 3700 m/sec in bone. If an incorrect average velocity is chosen, B-scan images (as well as other ultrasonic images, such as color flow images based on Doppler processing) develop image range and scan registration errors.
Under the assumption of a uniform tissue medium having constant sound propagation velocity, the presence of inhomogeneous tissues can result in image artifacts, range shifts, geometric distortions, etc. which degrade the ideal diffraction-limited lateral resolution, and increase the side lobes (which reduces the signal-to-noise ratio in the image).
These adverse effects of inhomogeneous and nonuniform tissue layers result in unknown phase aberrations associated with the inhomogeneities introduced across the transducer aperture. Many attempts have been made to overcome these aberrations using various signal processing techniques, generally referred to herein as phase aberration correction (PAC) processing.
PAC methods rely on comparison of the signal from one element or group of elements of an array (a correction element or group) to the signal received from another part of the array (a reference element or group), to develop a time delay for beamforming, of the correction group or element and the reference group. This time delay is optimized by any one of several methods, such as cross-correlation or speckle brightness. Such techniques are the subject matter of, for example, U.S. Pat. No. 4,852,577 (illustrative of investigations performed by Trahey et al. at Duke University) and U.S. Pat. No. 5,172,343 (illustrative of investigations performed by O'Donnell at General Electric), both incorporated herein by reference, to mention just two known PAC methods for ultrasound imaging systems.
After optimization, the value of the time delay for the correction element is fixed and then becomes part of the next reference group for optimizing a next correction element or group which is typically adjacent to the new reference group. The goal is to provide small adjustments in the time delay (the adjustments typically referred to as the "phase aberration correction profile") to correct for the defocusing effect of the forenoted tissue inhomogeneities.
Most array type ultrasound imagers use a 1-dimensional segmentation of the array (1D array) in the lateral dimension corresponding to the plane of the image. For good imaging performance, the elements are small and finely spaced in the lateral dimension (approx. 1 wavelength or less), however, the elevation (out of plane) dimension is fixed at a relatively large dimension (15 to 20 wavelengths, for example). This provides a fixed focal arrangement in the elevation dimension which defines the slice thickness of the planar image that is made.
As previously described, when the ultrasound signal propagates through an inhomogeneous medium such as the human body, variations in the index of refraction of the various tissues produce distortions of the wavefronts in both the lateral and elevational dimensions. The PAC process attempts to correct for such distortions in order to improve image quality, particularly contrast resolution. With a 1D array, it is obvious that the